1. Field of the Invention
The present invention pertains to tomography scanners and, more particularly, to positron emission tomography (PET) scanners designed for imaging small animals or humans.
2. Brief Discussion of the Related Art
Small animal PET scanners are commonly used in research facilities and, desirably, have high spatial resolution and uniformity and high sensitivity as described in “Molecular Imaging of Small Animals with Dedicated PET Tomographs,” Chatziioannou, Arion F., European Journal of Nuclear Medicine, Vol. 29, No. 1, January 2002. Commercial examples of such small animal PET scanners are the Concorde R4 and P4 “microPET” small animal PET scanners described in “Performance Evaluation of the MicroPET R4 PET Scanner for Rodents,” Knoess, Christof; Seigel, Stefan; Smith, Anne; Newport, Danny; Richerzhagen, Norbert; Winkeler, Alexandra; Jacobs, Andreas; Goble, Rhonda N.; Graf, Rudolph; Wienhard, Klaus; and Heiss, Wolf-Dieter, European Journal of Nuclear Medicine and Molecular Imaging, Vol. 30, No. 5, May 2003 and “Performance Evaluation of the MicroPET P4: A PET System Dedicated to Animal Imaging,” Tai, Y. C.; Chatziioannou, A.; Seigel, S.; Young, J.; Newport, D., Goble, R. N.; Nutt, R. E.; and Cherry, S. R., Physics in Medicine and Biology, 46 (2001) 1845-1862, and the Philips “Mosaic” small animal PET scanner described in “Design Evaluation of A-PET: A High Sensitivity Animal PET Camera,” Surti, S.; Karp, J. S.; Perkins, A. E.; Freifelder, R.; and Muehllehner, G., IEEE Transactions on Nuclear Science, Vol. 50, No. 5, October 2003. These scanners utilize dense arrays of small, individual detector elements, such as scintillation crystals ultimately viewed by photomultiplier tubes that encode the location of a scintillation event. The terms “detector elements,” “crystals” and “scintillating materials” are used interchangeably herein; however, it should be understood that the term “detector elements” includes any elements capable of detecting any type of radiation. The arrays are cylindrically arranged around a small diameter circle or polygon to form a mass of scintillating material that is nearly continuous in both the axial and circumferential directions. “Continuous” in this sense means that the individual scintillation crystals are as close together as possible such that any space between crystals is minimal and small compared to the crystal width and that the crystal positioning is replicated along the entire axial length of the scanning volume without appreciable or well defined gaps between rings of scintillation crystals. “Cylindrical detector array” as used herein means any geometric arrangement in which scintillation crystals or other detector elements circumferentially surround an imaging volume, e.g. in a circle, a polygon, an oval or the like, and have some axial extent. Prior art instrumentation for 3D PET imaging has focused on creating continuous axial and circumferential arrays of scintillator or other materials able to detect positron annihilation radiation emanating from a stationary imaging target or body, e.g. a small laboratory animal or a human. From these detected events, transverse section images of the radioactivity distribution from the body can be reconstructed that span the axial field of view of the device. The perceived need for continuity in detector arrays has been sufficiently compelling that prior art scanners have been specifically designed to avoid axially discontinuous arrays and have attempted to exploit novel array assembly methods, primarily optical/mechanical, that allow fabrication of continuous arrays of scintillation crystals, e.g. use of light guide coupling between crystal arrays and photodetectors, that allow close packing of crystals in both the axial and circumferential directions.
Dedicated PET scanners now on the market for either human or animal imaging targets use continuous cylindrical arrays of small scintillation crystals to define the imaging volume of the scanner. Discontinuous detector arrays e.g. paired, opposed flat panel detectors in time coincidence, are either mechanically rotated around the imaging target or the imaging target is rotated between fixed detectors to achieve the same result. See “A Rotating PET Scanner Using BGO Block Detectors: Design, Performance and Applications,” Townsend, David W.; Wensveen, Martin; Byars, Larry G.; Geissbuhler, Antoine; Tochon-Danguy, Henri J.; Christin, Anne; Defrise, Michael; Bailey, Dale L.; Grootoonk, Sylke; Donath, Alfred; and Nutt, Ronald, Journal of Nuclear Medicine, 1993; 34: 1367-1376, “Design and Characterization of IndyPET-II: A High Resolution, High Sensitivity Dedicated Research Scanner,” Rouze, Ned C. and Hutchins, Gary D., IEEE Transactions on Nuclear Science, Vol. 50, No. 5, October 2003, and “ECAT ART—A Continuously Rotating PET Camera: Performance Characteristics, Initial Clinical Studies, and Installation Considerations in a Nuclear Medicine Department,” Bailey, Dale L.; Young, Helen; Bloomfield, Peter M.; Meikle, Steven R.; Glass, Daphne; Meyers, Melvyn J.; Spinks, Terence J.; Watson, Charles C.; Luk, Paul; Peters, A. Michael; and Jones, Terry, European Journal of Nuclear Medicine, Vol. 24, No. 1, January 1997. The design of such scanners is commonly driven by the perceived need to create continuous, or virtually continuous, crystal arrays in the sense defined previously. For example, some of such scanners use individual light guides to connect crystals in an array to a phototube to eliminate the effect of “dead space” at the edges of phototubes. In other scanners, a continuous annulus of glass serves as a light guide to connect the cylindrical array of closely spaced small crystals to a bank of phototubes. In other scanners, such as the scanner disclosed in U.S. Pat. No. 6,288,399 to Andreaco et al, a large, axially continuous polygonal crystal array is created by centering, and packaging together, many small crystal arrays on clusters of four phototubes, a geometry that allows large arrays to be made by replication of this pattern as described in “The ECAT HRRT: Performance and First Clinical Application of the New High Resolution Research Tomograph,” Weinhard, K.; Schmand, M.; Casey, M. E.; Baker, K.; Bao, J.; Eriksson, L.; Jones, W. F.; Knoess, C.; Lenox, M.; Lercher, M.; Luk, P.; Michel, C.; Reed, J. H.; Richerzhagen, N.; Treffert, J.; Vollmar, S.; Young, J. W.; Heiss, W. D.; and Nutt, R., IEEE Transactions on Nuclear Science, Vol. 49, No. 1, February 2002. In each of these cases, technical innovations of one kind or the other are applied to allow small scintillation crystals to be packed closely together and to create detector arrays that are essentially continuous in both the circumferential and axial directions.
There are two primary reasons why the use of continuous arrays is deemed important. First, the idea that continuous arrays will intercept the largest fraction of annihilation radiation emanating from the target subject and, hence, will yield the maximum sensitivity for a particular ring diameter and axial length. Second, “classical” information theory has been thought to require continuous, regular and dense sampling of an imaging volume if imaging performance is to be as good as the system geometry permits. It has been generally believed that images reconstructed without dense and uniform sampling, i.e. without a continuous axial and circumferential distribution of scintillating material, would be of inferior quality, would contain artifacts or both. Degradation has been expected to increase if there were actual gaps in the detector array in either the circumferential or axial directions. In particular, while the effect on image quality of small circumferential gaps in detector arrays has been studied in some detail, see “Statistical Image Reconstruction in PET with Compensation for Missing Data,” Kinahan, P. E.; Fessler, J. A.; and Karp, J. S., IEEE Transactions on Nuclear Science, Vol. 44, No. 4, August 1997, “Correction Methods for Missing Data in Sinograms of the HRRT PET Scanner,” de Jong, Hugo W. A. M.; Boellaard, Ronald; Knoess, Christof; Lenox, Mark; Michel, Christiaan; Casey, Michael; and Lammertsma, Adriaan A., IEEE Transactions on Nuclear Science, Vol. 50, No. 5, October 2003, and “A Study of Image Errors Due to Detector Gaps Using OS-EM Reconstructions,” Yu, D.-C. and Chang, W., IEEE 1998, the literature contains little information about changes in image quality if a detector array possesses axial gaps, see “Design Optimization of the PMT-ClearPET Prototypes Based on Simulation Studies with GEANT3,” Heinrichs, U.; Pietrzyk, U.; and Ziemons, K., IEEE Transactions on Nuclear Science, Vol. 50, No. 5, October 2003.
While a continuous cylindrical array of scintillation crystals surrounding an imaging target is an effective way to intercept annihilation radiation from an imaging target, prior art methods possess significant practical disadvantages. For example, the need to connect individual or small groups of scintillation crystals to a phototube with a light guide adds complexity to the manufacturing process. More importantly, there is a demonstrable loss of scintillation light when the light passes into and through a light guide, thus, potentially reducing imaging performance. A similar loss occurs when a bulk light guide is used for the same purpose. That is, it has been believed that the “dead” regions at the edges of most photonic devices cannot be tolerated.
A number of three-dimensional (3D) image reconstruction methods have been proposed in recent years including the Fourier re-binning method (FORE) combined with some form of 2D image reconstruction, e.g. filtered backprojection (FBP) and the 3D re-projection method (3DRP), as described in “Exact and Approximate Rebinning Algorithms for 3-D PET Data,” Defrise, Michael; Kinahan, P. E.; Townsend, D. W.; Michel, C.; Sibomana, M.; and Newport, D. F., IEEE Transactions on Medical Imaging, Vol. 16, No. 2, April 1997 and “Performance of the Fourier Rebinning Algorithm for PET with Large Acceptance Angles,” Matej, Samuel; Karp, Joel S.; Lewitt, Robert M.; and Becher, Amir, Phys. Med. Biol. 43 (1998) 787-795, which are incorporated herein by reference. Iterative, statistical methods, such as 3D ordered subset expectation maximization (3D OSEM), as described in “Accelerated Image Reconstruction Using Ordered Subsets of Projection Data,” Hudson, H. Malcolm and Larkin, Richard S., IEEE Transactions on Medical Imaging, Vol. 13, No. 4, December 1994, which is incorporated herein by reference and 3D maximum a posteriori image reconstruction (3D MAP) as described in “High Resolution 3D Bayesian Image Reconstruction Using the MicroPET Small Animal Scanner,” Qi, Jinyi; Leahy, Richard M.; Cherry, Simon R.; Chatziioannou, Arion; and Farquhart, Thomas H., Phys. Med. Biol., 43, (1998) 1001-1013, both with system modeling, have also been introduced. A number of other algorithms that exploit the expectation maximization-maximum likelihood (EM-ML) approach with system modeling have also been studied as described in “Fast EM-Like Methods for Maximum “A Posterori” Estimates in Emission Tomography,” De Pierro, Alvaro R. and Yamagishi, Michel Eduardo Beleza, IEEE Transactions on Medical Imaging, Vol. 30, No. 4, April 2001. Each of these methods allows the 3D information potentially available in cylindrical PET scanners without collimators to be reconstructed into 2D slices that fully exploit the increased sensitivity associated with 3D data collections compared to purely 2D collections. To date, these methods have been used only for image reconstruction from scanners having continuous cylindrical arrays of scintillation crystals.